The field of the invention is nuclear magnetic resonance imaging methods and systems. More particularly, the invention relates to the measurement of blood flow using magnetic resonance imaging.
Any nucleus which possesses a magnetic moment attempts to align itself with the direction of the magnetic field in which it is located. In doing so, however, the nucleus precesses around this direction at a characteristic angular frequency (Larmor frequency) which is dependent on the strength of the magnetic field and on the properties of the specific nuclear species (the magnetogyric constant .gamma. of the nucleus). Nuclei which exhibit this phenomena are referred to herein as "spins".
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B.sub.0), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. A net magnetic moment M.sub.z is produced in the direction of the polarizing field, but the randomly oriented magnetic components in the perpendicular, or transverse, plane (x-y plane) cancel one another. If, however, the substance, or tissue, is subjected to a magnetic field (excitation field B.sub.1) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or "tipped", into the x-y plane to produce a net transverse magnetic moment M.sub.t, which is rotating, or spinning, in the xy plane at the Larmor frequency. The practical value of this phenomenon resides in the signal which is emitted by the excited spins after the excitation signal B.sub.1 is terminated. There are a wide variety of measurement sequences in which this nuclear magnetic resonance ("NMR") phenomena is exploited.
When utilizing NMR to produce images, a technique is employed to obtain NMR signals from specific locations in the subject. Typically, the region which is to be imaged (region of interest) is scanned by a sequence of NMR measurement cycles which vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques. To perform such a scan, it is, of course, necessary to elicit NMR signals from specific locations in the subject. This is accomplished by employing magnetic fields (G.sub.x, G.sub.y, and G.sub.z) which have the same direction as the polarizing field B.sub.0, but which have a gradient along the respective x, y and z axes. By controlling the strength of these gradients during each NMR cycle, the spatial distribution of spin excitation can be controlled and the location of the resulting NMR signals can be identified.
There are a number of well known NMR techniques for measuring the motion, or flow of spins within the region of interest. These include the "time-of-flight" method in which a bolus of spins is excited as it flows past a specific upstream location and the state of the resulting transverse magnetization is examined at a downstream location to determine the velocity of the bolus. This method has been used for many years to measure flow in pipes, and in more recent years it has been used to measure blood flow in human limbs. Examples of this method are disclosed in U.S. Pat. Nos. 3,559,044; 3,191,119; 3,419,793 and 4,777,957.
A second flow measurement technique is the inflow/outflow method in which the spins in a single, localized volume or slice are excited and the change in the resulting transverse magnetization is examined a short time later to measure the effects of excited spins that have flowed out of the volume or slice, and the effects of differently excited spins that have flowed into the volume or slice. Examples of this method are described in U.S. Pat. Nos. 4,574,239; 4,532,474 and 4,516,582.
A third technique for measuring motion flow relies upon the fact that an NMR signal produced by spins flowing through a magnetic field gradient experiences a phase shift which is proportional to velocity. This is referred to in the art as the "phase modulation" technique. For flow that has a roughly constant velocity during the measurement cycle the change in phase of the NMR signal is given as follows: EQU .DELTA..phi.=.gamma.M.sub.1 v
where M.sub.1 is the first moment of the magnetic field gradient, .gamma. is the gyromagnetic ratio and v is the velocity of the spins. To eliminate errors in this measurement due to phase shifts caused by other sources, it is common practice to perform the measurement at least twice with different magnetic field gradient moments as described in U.S. Pat. No. 4,609,872. The difference in phase at any location between the two measurements is then as follows: EQU .DELTA..phi.=.gamma..DELTA.M.sub.1 v
By performing two complete scans with different magnetic field gradient first moments and subtracting the measured phases in the reconstructed image at each location in the acquired data arrays, a phase map is produced.
The phase (.phi.) of each pixel in the phase map is determined by taking the ratio of the imaginary and real components of the NMR signal and then taking the arctangent. This phase map can be converted into a velocity image using the known flow encoding gradient first moment M.sub.1 and the constant .gamma.. The product of the velocity values in this image times the voxel size integrated over a blood vessel cross section provides an estimate of total blood flow through the vessel. The availability of such quantitative blood flow measurements in conjunction with the anatomical information provided by the magnetic resonance angiogram provides a more accurate method of assessing lesion significance. When applied to human coronary arteries, for example, the ratio of blood flow rates before and after the introduction of a pharmaceutical vasodilator provides an indication of coronary flow reserve which is a more accurate measure of lesion significance than morphological angiographic appearance.
Unfortunately, the phase difference method of measuring blood flow is not accurate when applied to small vessels, such as coronary vessels. The method assumes that the voxel is completely filled with spins moving at the measured velocity. When flowing spins occupy, for example, only one half the pixel volume and there are no static spins in the volume, the same phase value .phi. will be measured for both voxels, but the flow in one voxel is only one half that of the flow in the other voxel. This difference is shown in FIGS. 4A and 4B, where the vector 10 is the measured NMR signal for the full voxel and the vector 11 is the NMR signal for a half-full voxel. As shown in FIG. 4C, on the other hand, when a voxel is occupied by moving spins and static spins, the measured phase angle .phi.' will be altered by the signal component due to static spins indicated by vector 12. Rather than a true flow measurement as indicated by vector 13, therefore, an erroneous apparent flow indicated by vector 14 will be measured.
As shown in FIG. 4D, for example, where the blood vessels has a diameter only 2 to 3 times the voxel size, the phase difference flow measurement may be accurate for only one voxel 15 out of the nine voxels that indicate flowing spins in the vessel 16. As a result, when the indicated flow of all nine voxels are summed to measure the total flow in vessel 16, the error can be considerable. The error caused by the presence of static spins in the same voxel with moving spins can be corrected to a certain extend by using the phase difference method described by David M. Weber et al in Magnetic Resonance In Medicine, MRM 29:216-225 (1993). However, this technique does not solve the flow overestimation problem caused by the partial occupation of the voxel by non-resonant nuclear species in tissues surrounding the blood vessel.